Tri modal spectroscopic imaging

ABSTRACT

The present invention relates to a spectroscopic imaging system using autofluorescence and reflectance images to diagnose tissue. A preferred embodiment of the invention uses a plurality of light sources to illuminate a tissue region to provide the fluorescence and reflectance images, respectively.

CROSS REFERENCE TO RELATED APPLICATION

This application claims the priority of U.S. Provisional Application No.60/702,246, filed Jul. 25, 2005 entitled, TRI MODAL SPECTROSCOPICIMAGING. The entire content of the above application is beingincorporated herein by reference.

BACKGROUND OF THE INVENTION

Cancer is the second leading cause of death in the world. Each year,cancer kills over 500 thousand people in the United States alone(National Cancer Institute). Current cancer diagnosis methods usuallyinvolve two medical procedures. The first procedure is a wide-areasurveillance over the tissue, for example: mammogram, colposcopy,palpation, or visual examination. When warning signs are present, biopsyis performed on the suspicious tissue sites. However, many forms ofprecancerous and early cancerous lesions are difficult to detect usingthese traditional surveillance procedures. Therefore, there is a needfor wide-area surveillance systems capable of precancer detection.

Medical imaging modalities such as mammography and colposcopy haveproven vitally important for cancer diagnosis. So far, the majority ofimaging modalities focus on tissue structure or anatomy, which is notsufficient for detecting precancers at their earliest stages.Biochemical and subcellular morphological changes have been shown toaccompany precancer development. Thus, it is most beneficial to developnew cancer imaging modalities that provide tissue biochemical andmorphological information. Several new optical imaging modalities showgreat promise:

Confocal microscopy eliminates multiple scattering in turbid samples,producing thin section images with high resolution and contrast. Theimages produced are due to light scattered backwards at interfaces ofdifferent refractive index. Multiple scattered light is rejected bymeans of a pinhole, which selects only light traveling in straight-linepaths. The location and size of the pinhole, among other variables,determine the depth and lateral resolution of the system.

Optical coherence tomography (OCT) utilizes the coherence properties oflight to obtain cross sectional images of scattering media such asliving tissue. This technique employs low coherence light (i.e. lightwith a short coherence length) in a Michelson interferometer. Thespecimen is placed at the end of the sample arm. Back-scattered light iscombined with light returning from the mirror in the reference arm.Constructive interference occurs only when the distance to a scatteringinterface in the sample matches that to the reference mirror to withinthe coherence length. Depth is probed by scanning the reference mirrorposition and detecting the envelope of the interference signal.Cross-sectional images can be built up from multiple axial scans atdifferent transverse positions in the sample. As with confocalmicroscopy, image formation is again due to refractive index change.

Several groups have used polarized light to image superficial tissuesincluding using polarized light to enhance contrast in skin images byseparating the specular and multiple-scattered components of lightemerging from the skin surface or polarized gating can enhance theimages of surface and sub-surface structures in biological tissues.

Fluorescence is induced by the excitation of fluorophores in the tissue,usually with blue or ultraviolet (UV) light. Therefore, fluorescencecontains information about fluorophore concentration in the tissue.Two-photon microscopy (TPM) is capable of imaging fluorophores deepwithin a tissue sample. Tissue auto-fluorescence has also been used todetect neoplastic growths in-vivo.

Medical imaging modalities for precancer diagnosis can also employspectroscopy. Fluorescence spectroscopy imaging systems have been usedfor detecting cervical intraepithelial neoplasia and combinedfluorescence and reflectance spectroscopy methods are complementary forcancer diagnosis, making the use of the two techniques together morediagnostic than the use of either method separately.

SUMMARY OF THE INVENTION

Tri modal spectroscopy (TMS) can combine spectroscopic techniques togain biochemical, structural, and morphological informationsimultaneously. The present invention uses both fluorescence andreflectance imaging systems and methods for both in vivo and ex-vivomeasurements. Intrinsic fluorescence spectroscopy (IFS) is used, forexample, to obtain relative concentrations of fluorophores (e.g. NADHand collagen). Diffuse reflectance spectroscopy (DRS) providesinformation about the morphology and biochemistry of the stromal tissueand values of the absorption and reduced scattering coefficients,μ_(a)(λ) and μ_(s)(λ). Light scattering spectroscopy (LSS) determinesnuclear size, density, and distribution.

TMS has been implemented previously into a single point clinicalinstrument, to perform early cancer detection in-vivo. Three organ typeshave been measured including the esophagus, cervix, and oral cavity. Theresults demonstrated that TMS offered higher sensitivity and specificitythan any one spectroscopic technique alone.

Ex-vivo tissue measurements performed with LSS imaging system andmeasurements conducted using tri-modal spectroscopy implemented in asingle point instrument showed LSS imaging and TMS can be used forprecancer diagnosis. Implementing TMS into an imaging system providesthe advantage for screening larger regions of the body at faster speeds.The TMS imaging instrument can improve the sensitivity and specificityof cancer diagnosis with wide-area imaging systems while maintainingtheir abilities to provide real-time, non-invasive diagnosis.

Note that propagation of diffusely scattered light renders localizeddiagnosis more difficult because reflectance detected in one area of thetissue may carry contributions from other areas. Another challenge is toisolate single backscattering, diffuse reflectance, and intrinsicfluorescence so each component can be analyzed separately. The methodsseparate single scattering from diffuse reflectance and discriminate thelight scattering spectrum of certain size scatterers from that ofothers.

A preferred embodiment of the invention uses angular gating for lightscattering spectroscopy. This embodiment utilizes the measurement of atleast two reflectance spectra at two azimuthal angles to characterizetissue.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an LSS imaging system;

FIG. 2 is an image of color tissue sample;

FIG. 3 is a LSS imaging system;

FIG. 4 is data from a system;

FIG. 5 is a TMS imaging system;

FIG. 6 is a graphical illustration of data collected on the system ofFIG. 5;

FIG. 7 is an illustration of scattering in tissue

FIGS. 8 a and 8 b are measured spectra;

FIGS. 9 a and 9 b are angular scattering maps;

FIGS. 10 a and 10 b show spectral intensity variation for 10 and 1micron spheres;

FIGS. 11 a and 11 b are backscattering measurements;

FIG. 12 is a further preferred embodiment of an imaging system; and

FIG. 13 is another preferred embodiment of an imaging system.

FIG. 14 illustrates a scanning spot pattern used for imaging a region ofinterest.

FIG. 15 illustrates a portable system for use with a preferredembodiment of the invention.

FIG. 16 shows a schematic of a system for use in a handheld probe.

DETAILED DESCRIPTION OF THE INVENTION

A spectroscopic imaging system in accordance with a preferred embodimentof the invention can be for precancer detection over a relatively large1.3 cm by 1.3 cm area, for example. FIG. 1 depicts such a system 10. A75 W xenon arc lamp light source 12 (Oriel, Inc.) illuminates the tissue20. The light is collimated with half angle 0.5°, polarized 14, andtransmitted through one of 11 narrow-band (4 nm) filters 16 (EdmundScientific) to select the wavelength in the range 450-700 nm. Thereflectance from the tissue is diverted by the beamsplitter 18 to a 4fimaging system 22 that images the illuminated surface of the sample 1:1onto the CCD (Princeton Instruments, Inc.). The CCD consists of 512×512pixel array with pixel size 25×25 microns. A polarizer 24 along thecollection path selects the polarization state of the collected light.An iris 26 positioned at the center of the 4f lens system collects onlylight scattered into a solid angle corresponding to a half angle of0.5°. For each sample, two different reflectance spectra are measured,one with collection polarization parallel to that of excitation, and theother with perpendicular collection polarization. The polarizationgating technique is used to extract the single backscattering spectrumby subtracting perpendicular spectrum from the parallel spectrum. TABLE1 Comparison of the values of the mean nuclear diameters and standarddeviations of nuclear sizes in the colon adenoma of FIG. 2 andsurrounding non-dysplastic epithelium measured with LSS and usingstandard morphometry of the stained tissues. Morphometry LSS Normal Meannuclear diameter and 5.60 0.20 5.70 0.13 mucosa standard error ofmeasurement, μm Standard deviation of 1.01 0.82 nuclear diameters, μmAdenoma Mean nuclear diameter and 7.44 0.23 7.67 0.40 standard error ofmeasurement, μm Standard deviation of 1.59 2.19 nuclear diameters, μm

Experiments were conducted on physical tissue models of 15 microsphereson a diffusive medium, monolayers of T84 tumor colon cells, and ex-vivocolon tissue. From experiments with physical models and cells, theaccuracy of measuring size and relative index of refraction are 25 nmand 0.001 respectively were established.

LSS measurements of cancerous ex-vivo colon tissue were used to testnuclear sizing. FIG. 2 shows a color-coded image indicating differentpercentages of enlarged nuclei (greater than 10 μm in diameter). Thelight scattering measurements were compared with morphometrymeasurements done under a microscope. The results are compared onTable 1. Adenomas demonstrate clear nuclear enlargement over normalmucosa and size distribution is also more varied. Nuclear enlargementand variation are strong indicators of dysplasia. LSS results agree wellwith those of microscopy and wide area imaging has been demonstrated towork very effectively.

Using a plurality of excitation and emission wavelengths (fastEEM)measurements for non-invasive detection of dysplasia in three organs:uterine cervix, esophagus and oral cavity. TABLE 2 Performance ofdifferent spectroscopic techniques for separating SILs from non-SILsBiopsied non-SILs* vs SILs (%) Non-SILs† vs SILs (%) TechniqueSensitivity Specificity Sensitivity Specificity IFS 62 67 62 92 DRS 6957 62 82 LSS 77 71 77 83 TMS 92 71 92 90*Biopsied non-SILs include 21 colposcopically abnormal biopsied sitesthat were classified as MSE (5/21 sites) or SQM (16/21) sites.†Non-SILs in this case include 50 colposcopically normal sites and 21biopsied sites that were classified as SQM or MSETable 2 presents results from an extensive in-vivo study of uterinecervix dysplasia using the fastEEM. IFS, DRS, and LSS alone separatedbiopsied squamous intraepithelial lesions (SIL) from the biopsiednon-SILs and the biopsied SILs from all spectroscopically examinednon-SILs with sensitivities and specificities shown in the first 3 rows.TMS performed the same separations with sensitivities and specificitiesshown on row 4. The results show TMS offers better sensitivity andspecificity than any one spectral modality alone.

FIG. 3 is a system for LSS imaging 50. The light source 52 is a 20 HzQ-switched ND:YAG Laser (Opotek) pumped optical parametric oscillator(OPO). Light is coupled into a fiber (Fiberguide Industries, d=1 mm,NA=0.12) by a focusing lens (f=25 mm) and transmitted to the collimatinglens. (f=14 mm). The excitation is expanded 1:4 onto the sample 54through a 4f system (L₁ and L₂). The iris on the excitation path limitsthe excitation collimation to half-angle of 0.5 degrees. Polarizer Pdefines the excitation polarization. Light scattered at θ=180° from thesample travels along the excitation path prior to being redirected bythe beam splitter and imaged through another 4f system (L₂ and L₃) ontothe 12-bit CCD (Roper Scientific CoolSNAP HQ, 1392×1040 pixel array,pixel size 6.5×6.5 microns). The image is demagnified 5:1 by the optics.The iris on the collection path restricts the angular span of thebackscattered light collected to ±0.5 degrees. A calcite crystal 56splits the collected light into parallel and perpendicular polarizationsfor polarization gating [Backman 1999].

Control and data acquisition have been automated using softwaredeveloped with National Instruments LabView 6.1. Each set of spectraldata is acquired with excitation wavelength stepped from 470 nm to 670nm in 2 nm increments. The CCD 58 can be divided into 30 pixel by 30pixel areas. Each area corresponds to ˜1 mm² on the sample. Measuredspectra from each area, I_(m,∥)(λ₁, x,y) (parallel) and I_(m,⊥)(λ₁,x,y)(perpendicular) for (i=1, 2 . . . 100, 101), are the averaged spectraover all pixels in that area. One image is acquired for each excitationwavelength. Only one run is required to measure both parallel andperpendicular spectra because the calcite crystal divides thepolarization. This halves data-acquisition time. For each sample, wealso measure spectra with spectralon, I_(s,∥)(λ_(i),x,y) andI_(s,⊥)(λ_(i),x,y), for normalization. Normalization accounts forspatial and spectral variations of the system. The I_(m)'s arenormalized by the I_(s)'s and mean centered to one:${{\overset{\_}{I}}_{m, \parallel}\left( {\lambda_{i},x,y} \right)} = \frac{\frac{I_{m, \parallel}\left( {\lambda_{i},x,y} \right)}{I_{s, \parallel}\left( {\lambda_{i},x,y} \right)}}{{mean}\quad\left( \frac{I_{m, \parallel}\left( {\lambda_{i},x,y} \right)}{I_{s, \parallel}\left( {\lambda_{i},x,y} \right)} \right)}$${{\overset{\_}{I}}_{m,\bot}\left( {\lambda_{i},x,y} \right)} = \frac{\frac{I_{m,\bot}\left( {\lambda_{i},x,y} \right)}{I_{s,\bot}\left( {\lambda_{i},x,y} \right)}}{{mean}\quad\left( \frac{I_{m,\bot}\left( {\lambda_{i},x,y} \right)}{I_{s,\bot}\left( {\lambda_{i},x,y} \right)} \right)}$The single scattering spectrum is obtained with polarization gating:I _(SS)(λ_(i),x,y)= I _(m,∥)(λ_(i),x,y)− I _(m,⊥))(λ_(i),x,y)

Size information (mean and standard deviation) is extracted from theback scattering spectrum of each 1 mm² sample area. The extractionalgorithm compares all I _(SS)(λ_(i)) to Mie Theory calculations ofpolarization gating,${{\overset{\_}{I}}_{MIE}\left( {\lambda_{i},a,\delta} \right)} = {\frac{{I_{{MIE}, \parallel}\left( {\lambda_{i},a,\delta} \right)} - {I_{{MIE},\bot}\left( {\lambda_{i},a,\delta} \right)}}{{mean}\quad\left( {{I_{{MIE}, \parallel}\left( {\lambda_{i},a,\delta} \right)} - {I_{{MIE},\bot}\left( {\lambda_{i},a,\delta} \right)}} \right)}.}$This is the simulated single scattering spectra with mean centered toone. a is mean diameter and δ is standard deviation of size. Forinhomogeneous samples, one IMIE(λ_(i),a,δ) is computed for each area.The best least squares fit is found using fminsearch.m (Mathworks). Thesimulation program accounts for linewidth of the system, scatteringangle measured, and relative index of refraction (m=n_(m)/n₀) betweenthe sample (n_(m)) and the medium (n₀). Linewidth, scattering angle, andrelative index of refraction strongly influence the measuredbackscattering spectrum.

Systems and methods using a microscope cover slip sample containing 5 μmmicrospheres solution (Duke Scientific Corp.) accounted for the index ofrefraction spectral dependences of microspheres and water. FIG. 4 showsI _(SS)(λ_(i),x,y) (dotted line) from one area of the sample. Thesimulated results are the best fit I _(MIE)(λ_(i),a,δ) (solid line). Theagreement is excellent. Following the results obtained with 5 μmmicrospheres, we proceeded to validate the clinical prototype systemwith 5, 9, and 10 μm microspheres in water. Table 3 lists the measuredsize information (a, δ) for a randomly chosen area along side themanufacturer's specifications. The fits are excellent and the extractedparameters are within the manufacturing tolerances.

Table 3: Comparison of mean and standard deviation of diameter ofpolystyrene microspheres according to manufacturer specifications and asdetermined by the LSS imaging system 5 μm Spheres 9 μm Spheres 10 μmSpheres Mean Stdev Mean Stdev Mean Stdev. Manufacturer 5.01 0.035 8.9560.056 10.15 0.06 Spec. LSS Imaging 5.011 0.0204 8.921 0.036 10.093 0.026System

The schematic drawing of the system 100 for LSS including angular gatingis shown in FIG. 5. The light source 102 is a 300 W Xenon Arc Lamp(Oriel Thermo). A monochromator (Oriel Cornerstone) steps wavelength ofexcitation light with 10 nm FWHM spectral line width. The light is thencollimated by L1. A 2 cm diameter patterned mask 104 is imaged 1:1 on tothe sample by a 4-f system consisting of L2 and L3 with focal lengths of40 cm. P1 defines the polarization of the excitation. A long pass filterremoves light with wavelength below 385 nm. I1 limits the collimation ofthe excitation light to half angle of 0.5°. The reflectance from thesample 106 comes back along the path of the excitation before it isdeflected by a beam splitter 108 into the collection path. Another 4-fsystem (L4 and L5) with 1:5 demagnification images the surface of thesample onto the CCD detector 110 (Photometrics Sensys, 768×512 pixelarray, pixel size 9×9 microns). The iris 112 on the collection pathrestricts the angular span of collected backscattered light to ±0.5degrees.

The sample is illuminated by the excitation light with wavelengthranging from 450 nm to 700 nm in 5 nm increments. At each wavelength,the CCD acquires an image. Exposure time is varied for different sampletypes to utilize most of the 4096 counts available on each pixel. Aspectrum, I_(m)(λ_(i),x,y) for (i=1, 2 . . . 50, 51), is obtained foreach unit area, 20 pixel×20 pixel, on the CCD by plotting the recordedintensities at the different wavelength images. For each spectrum from asample we measure a corresponding spectrum is measured,I_(s)(λ_(i),x,y), for normalization purposes by replacing the samplewith a broadband dielectric mirror (Thorlabs, Inc.). This accounts forspectral and spatial variations of our system because reflectance fromthe mirror is above 99% over the entire spectrum. Measured reflectancespectra were obtained from a 1 mm thick, 1″ diameter solution of 10 μmmicrospheres (Duke Scientific Corp.) immersed in a density matchingfluid (80% water and 20% glycerol). The standard deviation of themicrospheres' diameter was 0.058 μm. The optical thickness of themicrospheres, λ, was about 0.2. Spectra were measured for backscatteringangles θ=178°, Φ=0° (I_(m,0)(λ_(i))) and Φ=90° (I_(m,90)(λ_(i))) FIG. 6shows two of the normalized spectra, I_(m,0)(λ_(i),x,y) andI_(m,90)(λ_(i),x,y) (solid lines). Simulated results (dotted lines) werecalculated using Mie theory.

Spatial gating is a powerful gating technique that separates lightscattered by superficial layers of the sample from those photons thathave traversed deeper sections.

Light reflected from the superficial layers have undergone fewscattering events, including those that have scattered only once (singlescattering intensity, I_(S)), while photons returning from deeper layershave scattered more (diffuse reflectance, I_(D)). FIG. 7 illustrates howspatial gating is used to separate single scattering from diffusereflectance. The photons experiencing single backscattering enter andleave the sample from the same region (black). The photons experiencingmultiple scatterings will enter and leave the sample either through thesame region (blue) or a different one (red). The blue photons undergoneither single scattering nor diffuse reflectance. Photons that undergomore scattering events are increasingly likely to exit from a differentregion. Therefore, when the sample is imaged onto the CCD, singlescattering will only appear in lighted areas while multiple scatteringwill appear in light and dark areas. Spatial gating uses a specificpatterned mask in the system of FIG. 5. The mask consists of twoparts: 1) a 1″ iris 114 whose opening controls the size of theexcitation and 2) a transparent glass with opaque regions providing apattern that is imaged by a 4f system (L2 and L3 in FIG. 5) onto thesample to create dark regions. To obtain majority single scattering,measure the intensity in a lighted area and subtract measurements fromthe closest dark area as illustrated below:Light area: I _(light) =I _(S) +I _(D)Dark area: I_(dark)=I_(D)Single scattering intensity: I _(light) −I _(dark)=I_(S)  (Equation 1)

Measurements with tissue phantoms validate spatial gating. The tissuephantoms are two-layer models designed to simulate epithelial cells ontop of underlying tissue. Epithelial cell nuclei are approximated by 10mucon microspheres (Duke Scientific Corp. and Polysciences Inc.) in thedensity-matching fluid used for system calibration. The optical densityis about 0.2. Beneath the microspheres solution is 10%; Intralipid.Intralipid models underlying tissue and it is diluted to have reducedscattering coefficient, μ_(s)′, measured with the procedure describedin, similar to that of biological tissue. The excitation mask used forthe experiments divides the excitation light into many unit areas(approximately 1 mm²), within each is embedded a small unilluminatedarea. This pattern is a “Dark Spot”. The LSS signal for a unit area isthe average spectrum of the lighted area minus that of the dark spot(Equation 1). FIG. 8 a demonstrates the effectiveness of spatial gating.The black solid line is the total reflectance spectrum measured from theilluminated area, I_(light). The reflectance spectrum from theneighboring local dark area (I_(dark), dotted black line) contains onlythe diffuse reflectance component. The difference between the twospectra, I_(S), is shown by the solid blue line. FIG. 8 b is a close upof I_(S) (solid line) and the best fit simulated result (dotted line).

Dark spot size is important because utilizing a smaller spot will reducethe unwanted blue photons of FIG. 7 and permit more lighted area, wheresingle scatterings occur. Spot size is limited by our system'sresolution, which is determined by the small collection numericalaperture (NA) intrinsic to angle dependent LSS measurements. This placesa lower limit on the dark spot size of roughly diameter=50 μm. For DRSand IFS measurements, we use a “Light Spot” pattern as the mask. LightSpot is the inverse of Dark Spot, in that each unit area has a smallilluminated region. This mimics a bundle of fastEEM contact-probes.

Measurements indicate spatial gating is a viable technique forseparating the light scattering and diffuse reflectance components ofthe total reflectance spectrum. This complements the mathematicalmodeling used by the TMS fastEEM and polarization gating.

Angular gating is a gating technique capable of favoring singlescattering from certain size scatterers over that from other sizes. Itis implemented by collecting single backscattering at certain anglescorresponding to favored sizes.

Angular gating exploits the non-isotropic scattering of large particles.The scattering intensity from a particle illuminated by a plane-wave isa function of scattering angle (θ), azimuthal angle (Φ), particle radius(a), light wavelength λ, and the relative index of refraction of theparticle with its surrounding medium (m). FIG. 9 shows angularscattering maps of two different size scatterers. The radial directioncorresponds to backscattering angle θ while the angular direction is Φ.Observe the 10 μm particle scatters in distinct lobes while the smaller1 μm scatters isotropically. FIG. 10 shows the spectral plots of FIG.9's Φ=0° and Φ=90° lobes. Note that the two 10 μm scattering spectra arevery different while the 1 μm ones are almost identical. When one takesthe difference between two spectra, the oscillatory feature is enhancedfor 10 μm particles but suppressed for 1 μm particles. Based on thisfinding, Size Discrimination Angular Gating (SDAG), which produces aspectrum by taking the difference between Φ=0° and Φ=90° measurements.Implement SDAG by measuring two spectra. First, we set the excitationpolarizer, P1 of FIG. 5, parallel to the surface of the table. P2 is setparallel to P1 and one full spectrum is recorded. Next, P1 and P2 arerotated 90° and another spectrum is recorded. Rotating the twopolarizers is equivalent to moving I2 between the Φ=0° and Φ=90°positions on the fourier plane of L4. The difference of these twospectra will enhance the scattering signal from particles with certainsizes, chosen by the scattering angle θ.SDAG: I _(SDAG)(λ_(i))=I(λ_(i),φ=0°)−I(λ_(i),φ=90°)To characterize SDAG, a microspheres sample with 10 μm, λ˜0.2, and 1 μm,λ˜2 immersed in the index matching solution used for calibration. Thesolution rests on top of an absorption neutral density filter.Scattering is measured at θ=178.1°, Φ=0° and Φ=90°. The peakbackscattering lobes for 10 μm scattering are at these angles. Averagedresults from the entire sample area are presented in FIG. 11 a (Solidcurves). Theoretical results (dotted curves) are computed using MieTheory with the manufacturer's specifications for the 10 μm spheresbecause Angular Gating at θ=178.1° should discriminate against othersizes. I_(SDAG)(λ_(i)) deviates little from the theoretical 10 μmspectrum despite a 10:1 optical density ratio between 1 μm and 10 μmmicrospheres. We also test the effectiveness of Angular Gating underincreasing optical density ratios. FIG. 11 b has ratio of 20:1, yet the10 μm frequency is visible in I_(SDAG)(λ_(i)).

Light scattering from tissue is composed of nuclei and smallerorganelles scattering. Part of current precancer diagnosis relies onfitting the light scattering spectrum to a size distribution and indexof refraction. Angular Gating favors certain size scatterers over othersin the LSS spectrum, potentially allowing more accurate sizedistribution and index of refraction extraction. The respectivescatterer sizes were used in the measurements because 10 μm approximatesnuclear diameter while 1 μm approximates mitochondria diameter, one ofthe more abundant smaller organelles. In squamous and columnarepithelial cells, the optical density ratio between mitochondria andnuclei is approximately 10:1, as estimated from. Measurements havedemonstrated successful size discrimination under similar conditions.Size Discrimination Angular Gating can significantly enhance the nuclearsignal in LSS.

A TMS imaging instrument is used for detecting cervical dysplasia, oralcavity dysplasia, and breast lesions. Images are collected bysimultaneously collecting image data with a two-dimensional spatialarray of pixel elements to collect light from a corresponding area of atissue surface. Preferably, the imaging detector has at least a 500×500array of pixel elements. Such an image of the region of interest iscollected for each of the fluorescence and reflectance images used inthe system. In a preferred embodiment separate light sources can be usedfor the fluorescence and reflectance images, the light from each sourcebeing delivered through a common light delivery system, with the imagesbeing collected through at least a partially common light path of thelight collection systems.

The system includes studying LSS, DRS, and IFS in the imaging mode, thespatial and angular gating techniques. Table 4 shows the requirementsfor LSS, DRS, and IFS measurements. To achieve these objectives, thesystem 200 is displayed in FIG. 12 For LSS data acquisition, awavelength tunable optical parametric oscillator 202 (OPO) is used toprovide single wavelength excitation from 400-700 nm. The OPO allows 2orders of magnitude increase in excitation power on the sample over theXenon Arc Lamp 204 due to its better collimation, which leads tostronger light scattering signals when exposure time is constrained. Thewavelength tuning speed of the OPO is also higher than that of themonochromator 206. This further expedites data collection and avoidscell or ex-vivo tissue degradation problems during measurements. Duringdata acquisition, light from either the OPO or the arc lamp is selectedto excite the multimirror 210 by a flippable mirror 208. The digitalmultimirror device (DMD, Texas Instruments) is a 1024 pixel×768 pixelarray of micromirrors which can be separately directed at two differentangles. Each mirror pixel is estimated to be 14 microns by 14 microns.The DMD acts as a spatial light modulator and is used to define preciseexcitation patterns. For LSS, a dark spot pattern of excited lightreflected by the multimirror is imaged and magnified 1:2 onto the sampleby L1 and L2. Polarizer P defines the excitation polarization. Thebeamsplitter and I1 of FIG. 5 are replaced by a mirror, <3 mm diameter,placed at the focal point of L1, L2, and L3. This mirror serves 2purposes: it defines the divergence of the excitation and redirects thelight towards the sample. With L2 focal length of 16 cm, the excitationdivergence is limited to a half angle of 0.5° which leads to spatialresolution of approximately 50 μm. Light scattered from the tissue 220not at scattering angle θ between 179° and 181° will miss the mirror 222during collection and be imaged onto the CCD 224 by L2 and L3. ApertureA around the mirror defines the ranges of θ's and Φ's permitted to passthrough to L3. The aperture is opened to permit ±0.5° about the desiredθ and Φ. For DRS data acquisition, a Xenon arc lamp and monochromatorare used to provide single wavelength excitation from 300-700 nm. Lightpropagates through the system as in LSS, except the multimirror reflectsa light spot pattern and aperture A is removed to collect as muchreflectance as the lenses will permit. For IFS data acquisition, the arclamp and monochromator provide single wavelength excitation from 300-400nm. Light propagates through the system as in DRS, except the narrowline width filters, F, are rotated with each image acquired to definecollected light wavelength.

Data acquisition and processing will be automated by a computer programwritten on Labview 7.0 (National Instruments). Algorithms are adaptedfrom those used in the fastEEM. TABLE 4 Requirements for LSS, DRS, andIFS measurements LSS DRS IFS Excitation 400-700 nm 300-700 nm 300-400 nmwavelength Excitation ±0.5 degrees N/A N/A collimation Mask type DarkSpot Light Spot Light Spot Polarized Yes no no Excitation requiredCollection ±0.5 degrees N/A N/A collimation

The imaging instrument is constructed from optimal implementations andspecifications tested on the TMS fastEEM. It is used for real timewide-area studies of uterine cervix, oral cavity, esophagus, and breast.

The system's 300 optics are arranged as illustrated in FIG. 13 to meetthe TMS requirements on Table 4. The optical setup and data acquisitionsoftware are similar to those of the prior system. The multimirror is nolonger used as the excitation mask. Instead, a dark spot pattern mask302 for LSS is made of glass with opaque dots. For DRS and IFS lightspot masks, light from the arc lamp 304 into the proximal end of a fiberbundle. The distal end is attached to the source plane to form the lightspot pattern. These changes improve transmission efficiency and reducecomplexity. A wheel of narrow linewidth filters replaces themonochromator because filters scan the spectrum much faster. The CCD 306is set to high sensitivity for IFS data acquisition because speed isimportant. The distance from lens G32-923 (Edmund Optics) to the sample308 is roughly 15 cm. The collection optics, with the aperture betweenlenses G32-923 and G32-327 fully opened, can have F number approximately4.

LSS measurements with a dark spot mask use more power. A typicalmegapixel CCD has a well depth of 100000 photoelectrons/pixel andquantum efficiency 50%. To maximize signal to noise ratio, there areapproximately 200000 photons/pixel on the CCD. For LSS measurements,scattered light reaches the CCD only if is within ±0.5 degrees of themeasured θ. Therefore, if we assume the tissue is a Lambertianscatterer, a good approximation for diffusive media like tissue, roughly2.5×10¹⁵ photons are needed to exit the tissue. This corresponds toabout 1 mJ of energy spread over the entire area (4 cm²) at everywavelength measured. The excitation light must be collimated to ±0.5degrees and have relatively narrow line width (˜10 nm). Between theexcitation mask carrying the dark spot pattern and the sample, the lightpasses through a linear polarizer P. This cuts energy by roughly 60%.Therefore, about 3 mJ is needed from monochromatic, collimated lightilluminating the mask. The OPO emits 30 mJ/pulse (5-8 ns pulse, 1 nmline width) in the visible regime and operates at 20 Hz. Assumingefficient light transmission from the laser to the mask via a fiber, onepulse provides enough power for one LSS image. DRS and IFS measurementscan be acquired with a Xenon arc-lamp. Based on the power requirementsof a fastEEM contact-probe, this light source provides ample energy forDRS. For IFS however, use the high sensitivity mode of the CCD is usedbecause IFS signals pass through a narrow line width filter. The OPOsystem, a 300 W Xenon arc lamp, and a PhotonMax CCD (PrincetonInstruments) meet the requirements.

The total data-acquisition time depends on the number of measurementimages required and light source wavelength scan speed. A preliminaryestimate can be based on spectral features. The key features of the DRSspectrum are the hemoglobin absorption dips and the slope. We estimate12 points placed about the Hb dips can adequately characterize thespectrum. A total of one DRS spectrum is required. The IFS spectra havepeaks at emission wavelengths corresponding to NADH, tryptophan,collagen, FAD, and porphyrin. An estimate 11 emission wavelengths cancharacterize the spectrum for each excitation wavelength. 3 excitationwavelengths at 308 nm, 340 nm, and 400 nm have high diagnostic value.For LSS spectra, 11 wavelengths were measured for each polarizationgating setting. However, the speed of data acquisition is limited by theOPO wavelength sweep speed, which is 3 seconds from 400 nm to 700 nm. Intotal, TMS imaging will take approximately 160 images. LSS data can beacquired in 6s. IFS can be acquired in 2s and DRS in 1s. The PhotonMaxhas frame rate >25/s, so image acquisition is not a limiting factor. Todecrease acquisition time, we will investigate laser systems that cannotemit at as many excitation wavelengths as the OPO, but can scan thespectrum from 400-700 nm faster because we do not require 60 images/LSSspectrum.

The TMS imaging system can be used for two purposes: (1) as a diagnostictool to detect, diagnose, characterize and (2) to guide biopsy ofdysplastic intraepithelial lesions. Because this instrument allows earlydysplastic transformation to be characterized without the need fortissue removal, it enables the natural history of such lesions to bestudied in vivo.

TMS imaging operates like a bundle of many TMS contact-probe systems forDRS and IFS. One fastEEM contact-probe can analyze roughly 1 mm² oftissue, so a bundle of several hundred can simultaneously analyze a fewcm² of tissue. TMS imaging illuminates the tissue with a number of lightspots and collects the reflectance from the area around each spot. Eachlight spot and the signals collected around it are similar to a fastEEMprobe. TMS imaging results are comparable to the successful results fromsingle point TMS, but there are differences in the comparison that needto be considered. A contact probe excites and collects reflectance fromthe tissue with larger acceptance angle than the imaging system. As aresult, the imaging system has better angular resolution, crucial forLSS, but requires more excitation power to collect equal reflectance.Spectral information cannot be acquired in one image because the CCD'spixels record spatial information. Therefore, multiple images, each at adifferent wavelength, are required. This increases acquisition time.When multiple light spots excite the tissue simultaneously, excitationfrom one spot can enter the detection area of another spot. This iscross talk, which hinders acquisition of local reflectance information.

To minimize unwanted multiple scatterings and enhance LSS signal, limitsof dark spot size are modified to reach the resolution limit, which wasroughly 50 μm. The effectiveness of different dark spot shape andspacing. The main challenge for optimizing the dark spots is designingaccurate and precise masks. For example, changing dark spot size, shape,or spacing all require etching a mask.

Size Discrimination Angular Gating can be evaluated by progressivelyincreasing 1 μm optical density and measuring clarity of the favored 10μm spectrum at θ=178°. The effects of size, angle and index ofrefraction distributions on technique effectiveness. SDAG measurementsare used with different scatterer sizes at their optimal angles and withvarying degrees of excitation collimation. The results of thesemeasurements will demonstrate the strengths and limitations of angulargating. Data acquisition time with the clinical system can be reduced ifrelaxing excitation collimation minimally influences results.

To effectively measure diffuse reflectance contribution from one regionof the sample, the “Dark Spot” excitation pattern requires modificationbecause the red photons of FIG. 7 detected in one area originated frommultiple origins. This is cross talk.

For a point source excitation on a homogeneous medium with scatteringand absorption coefficients μ_(s)′ and μ_(a), Farrell et. Al derived thediffuse reflectance flux perpendicular to the surface at differentdistances ρ from the source to be:${R(\rho)} = {\frac{z_{0}a^{\prime}}{4\quad\pi}\left\lbrack {{\left( {\mu_{eff} + \frac{1}{r_{1}}} \right)\frac{{\mathbb{e}}^{{- \mu_{eff}}r_{1}}}{r_{1}^{2}}} + {\left( {1 + {\frac{4}{3}A}} \right)\left( {\mu_{eff} + \frac{1}{r_{2}}} \right)\frac{{\mathbb{e}}^{{- \mu_{eff}}r_{2}}}{r_{2}^{2}}}} \right\rbrack}$The symbols are defined as:$r_{1} = \left\lbrack {z_{0}^{2} + \rho^{2}} \right\rbrack^{\frac{1}{2}}$$r_{2} = \left\lbrack {\left( {z_{0}\left( {1 + {\frac{4}{3}A}} \right)} \right)^{2} + \rho^{2}} \right\rbrack^{\frac{1}{2}}$$a^{\prime} = \frac{\mu_{s}^{\prime}}{\mu_{a} + \mu_{s}^{\prime}}$$\mu_{eff} = \sqrt{3\quad{\mu_{a}\left( {\mu_{a} + \mu_{s}^{\prime}} \right)}}$$z_{0} = \frac{1}{\mu_{a} + \mu_{s}^{\prime}}$

This assumes all scattering events are isotropic. Parameter A depends onthe refractive index of the medium and is roughly 3.2 in colon tissue.Assume the medium is human tissue and its entire surface is illuminatedby excitation light. With this assumption even though imaging systemshave finite illumination areas to achieve a closed-form solution. Thetotal diffuse reflectance flux perpendicular to the tissue is given by:$R = {{\int_{0}^{2\quad\pi}{{\mathbb{d}\theta}{\int_{0}^{\infty}{{R(\rho)}\rho{\mathbb{d}\rho}}}}} = {\left( {1 + {\mathbb{e}}^{{- \frac{4}{3}}A\quad\mu_{eff}z_{0}}} \right)\frac{{\mathbb{e}}^{{- \mu_{eff}}z_{0}}}{2}}}$

The desired tissue parameters μ_(s)′ and μ_(a), appear in μ_(eff)*z₀.The above equation can only be solved for μ_(eff)*z₀ and not the desiredparameters μ_(s)′ and μ_(a) separately. This is not a problem for singlepoint systems. To address cross talk, we use a form of spatial gatingwhere the excitation pattern is again divided into many unit areas, butinstead of dark spots, each area has a “Light Spot”. The local diffusereflectance spectrum is the average spectrum in the area. The size of aunit area is set such that diffuse reflectance recorded in the areacomes largely from the local light spot. The illuminated spot is made assmall as physically possible. IFS imaging may also encounter some crosstalk difficulties because it is extracted using the reflectancespectrum. By adjusting the size of the unit areas with tissue phantomscontaining scatterers and fluorophores to minimize residual cross talk.The influence of light spot shape and spacing can be addressed.

Non-contact DRS and IFS imaging collects reflectance with a smalleracceptance angle than single point systems. This immediately increasesdata-acquisition time, but may also influence spectral features.Different acceptance angles can influence reflectance and fluorescencespectra. By varying collection angle in the system and by adjusting thecollection aperture these can be addressed.

Gating techniques can be used with tissue phantoms and the system hasdemonstrated successful separation of LSS and DRS and effectivediscrimination of scatterer size.

The TMS imaging system is used to detect precancer and early cancer inthe cervical and oral cavities. The extracted tissue biochemical andmorphological information is compared to histology and correlated withpathology.

Cell rafts will have oral cells, grown from biopsied tissue, supportedby a layer of collagen. The raft is immersed in growth media up to thelevel of the cells. We will create rafts with normal or malignant cells.Cell rafts can be used for oral cancer imaging measurements on bothnormal and cancerous samples.

Within an improved fitting algorithm that minimizes the least squareerror between fit and LSS data with minimal assumptions for tissueparameter extraction. This forms the LSS portion of the real-time dataprocessing algorithm in the clinical system. The method of was sensitiveto noise because uncertainty often appeared as oscillatory features. Themethod of assumed a Gaussian distribution. The Mie Theory FittingAlgorithm, MTFA, will solve the optimization problem min∥Ax−b∥_(x≧0).Parameter A is a matrix of scattering intensities computed at differentsizes, wavelengths, angles, and index of refractions using Mie Theory.Parameter b is an experimentally recorded spectrum and x is the best fitsize distribution. Our assumptions are uniform scattering angle, indexof refraction, and x≧0. Unfortunately, non-linear constraints greatlyincrease computation time. MFTA is optimized to reject experimentalnoise and reduce computation time. DRS and IFS data processingalgorithms have been developed for the fastEEM and are adapted to theTMS imaging system.

The instrumentation and diagnosis algorithms analyze the data. Theorgans are uterine cervix, oral cavity, esophagus, and breast. Theimaging system is less invasive than the TMS fastEEM studies because nophysical contact is required.

With one button press, the first generation system will acquire LSS,DRS, and IFS measurements in under 10s. Short data acquisition timemakes it easier for the patient to remain motionless. The doctor willmark regions in the data acquisition area for biopsies. The results arespatially correlated with wide area spectroscopic measurements. Thesystem provides a precancer diagnostic tool for cervix and oralcavities. The user presses a button to begin data acquisition, whichfinishes within 2 seconds. Then, rapid data processing produces adiagnostic map that color codes precancer risk and displays on thecomputer screen. Therefore, the user can see in real time which areasare at greater risk for precancer. The method correlates spectroscopydiagnosis with pathology analysis and can guide biopsy or papa smears.

The imaging can be conducted in multiple shots. For each shot, fourcollinear light spots illuminate four of the diagnostic regions on thetissue surface and reflectance spectra from these regions are imagedonto the CCD via a spectrograph. The four light spots 402 are shown onFIG. 14. To cover the entire area 400, the line of spots is scannedacross the tissue surface.

A general layout of the TMS imaging system 450 for cervix and oralcavity studies, for example, is shown in FIG. 15, along with criticalrequirements. Module 1 houses a broadband white light source 452 and a340 nm nitrogen laser 454, along with a coupler 456 to couple light intofour delivery fibers 458. The fibers are bundled together to deliverlight from Module 1 to Module 2, which is a compact and light-weighthandheld unit containing the system optics. At the input end of Module2, the fibers' distal ends are aligned collinearly. A bundle of 8collection fibers 462 transfers light from Module 2 to Module 3, whichhouses the spectrograph 464 and CCD 468. Module 4 houses the computer468 and software for instrument control and data processing. Modules 1,3 and 4 will be placed on a mobile, compact cart suited for hospitalsettings. TABLE 5 Constraints for Module 2 Size No larger than a singlelens reflex camera. Weight Less than 5 lbs. Separation 15 cm to 30 cmvariable (maybe closer in the between oral cavity) instrument andpatient Sampling Four approximately 0.6 mm diameter white light spotsilluminate the tissue. These spots are arranged collinearly and spaced 5mm apart. Collection 2° ± 0.5° from the direct backscattering angledirection and at azimuthal angles Φ = 0° and Φ = 90°. Sweeping The lineof light spots must be swept across the patient to cover a 2 cm × 2 cmarea with resolution no larger than 1 mm × 1 mm. Each unit area yields adiagnosis. Light 340 nm excitation wavelength 370 nm-700 nm collection(Fluorescence) Light 370 nm-700 nm illumination wavelength 370 nm-700 nmcollection (Elastic scattering) Image <10 s. acquisition time Focus Dueto the uneven nature of the tissue surface, some regions of the 2 cm × 2cm area will be in focus while other regions will be out of focus. Focusshould be corrected either by moving components within the system(autofocus?) or by data processing after all data has been acquired.

TABLE 6 Additional constraints for combined Modules 2/3 CCD Quantum >50%from 370 nm-700 nm Efficiency Readout rate >10 frames/s Chip size >=8 mm× 8 mm Pixel number >=512 × 512 Full well >=100 ke⁻ Dark current <0.1e⁻/pixel/sec @ operating temperature Spectrograph Range 370 nm-700 nm onthe corresponding CCD Line width FWHM <=10 nm on the corresponding CCDEfficiency >50% over wavelength range Slit height >=1 cm Acceptance F/#<=4FIG. 16 illustrates a preferred embodiment of a system 500 of the insideof Module 2. This includes a handle 502 and a distal probe 504 forinsertion into body cavities. Light delivered from Module 1 is focusedby lens L1 (f=15 cm) onto a 5 mm diameter rod mirror. The chromaticaberration of L1 is sufficiently small in the wavelength range of thesystem, 340 nm-700 nm, such that the light is able to focus onto the rodmirror 508. L2 (f=30 cm) collimates the light onto mirror M0a. L2 mustalso experience minimal chromatic aberration from 340 nm-700 nm. M0a andM0b redirect the light to the patient. The focal lengths of L1 and L2can vary, but the pattern of light delivered from Module 1 must beimaged onto the tissue, which is a fixed 30 cm from M0b. M0a and M0b canbe adjusted to sweep the line of spots across the tissue surface. Lightreturning from the tissue travels through M0a, M0b and L2. Two solidangles of light centered at θ=2°, φ=0° and φ=90°, are permitted to passby the aperture 520 through to M1 and M2. The aperture is located onefocal length behind L2. The two cones of light passing through theaperture are redirected by M1 and M2 towards L3 (f=10 cm). M1 and M2separate the two beams slightly such that L3 will form two separateimages of the line of spots. Each of the 8 images of the original 4spots of light is coupled into a fiber that delivers light to Module 3.

In FIG. 16 M indicates mirrors, P indicates linear polarizers, and Lindicates lenses. The blue and red rays are the φ=0° and φ=90° raysrespectively. In this figure, M2 appears to obstruct the light path offM1. M2 and M1 are positioned in different planes such that noobstruction occurs. The aperture in the top right corner is a expandedview of the aperture in the system. The top hole permits the φ=0° ray topass and the right hole permits the φ=90° ray to pass.

Folded beams are used to reduce the size of the system to provide ahandheld probe. Modules 2 and 3 are integrated into a single unit whilerespecting size and weight requirements.

While the present invention has been described herein in conjunctionwith a preferred embodiment, a person with ordinary skill in the art,after reading the foregoing specification, can effect changes,substitutions of equivalents and other types of alterations to thesystem and method that are set forth herein. Each embodiment describedabove can also have included or incorporated therewith such variationsas disclosed in regard to any or all of the other embodiments. Thus, itis intended that protection granted by Letters Patent hereon be limitedin breadth only be definitions contained in the appended claims and anyequivalents thereof.

1. A spectroscopic imaging system comprising: a light source; an imagingdetector that detects fluorescent and reflectance data to provide animage of tissue; and a data processor that processes the image todiagnose the tissue.
 2. The system of claim 1 further comprising a maskand plurality of filters.
 3. The system of claim 1 further comprising amoveable mirror device.
 4. The system of claim 1 further comprising acomputer program to diagnose cancerous tissue.
 5. The system of claim 1wherein the light source comprises a collimating light source.
 6. Thesystem of claim 5 wherein the light source comprises a laser.
 7. Thesystem of claim 1 wherein the light comprises a broadband light source.8. The system of claim 1 wherein the light source comprises a tunablesource.
 9. The system of claim 3 wherein the movable mirror comprises adigital micromirror array.
 10. The system of claim 3 wherein the movablemirror switches between a first light source and a second light source.11. The system of claim 5 wherein the collimating light source is anoptical parametric oscillator.
 12. The system of claim 1 furthercomprising a spatial gate.
 13. The system of claim 1 further comprisingan angular gate.
 14. The system of claim 1 wherein the processordetermines a size of a cellular structure from the reflectance spectrum.15. The system of claim 1 further comprising a fiber optic device. 16.The system of claim 1 wherein excitation light for the fluorescence andreflectance images coupled to a single light delivery system.
 17. Thesystem of claim 1 further comprising a light collection system collectthe fluorescence and reflectance images along a common optical path. 18.The system of claim 1 wherein the detector simultaneously detects anarea of a region of interest with a two dimensional array of pixelelements.
 19. The system of claim 19 wherein the detector comprises atleast a 500×500 array of pixel elements that form an image of an area oftissue have at least 500×500 image pixels.
 20. The system of claim 1further comprising a collection optical system that collects a polarizedcomponent of light.
 21. The system of claim 1 further comprising apolarizer.
 22. The system of claim 1 further comprising an aperture thatseparates a polarization component.
 23. The system of claim 1 furthercomprising a handheld probe for insertion into a body cavity.
 24. Thesystem of claim 23 wherein the handheld probe is coupled to the lightsource with a fiber optic delivery cable and to the detector with afiber optic collection cable.
 25. The system of claim 23 wherein thehandheld probe comprises a handle and a distal probe section forinsertion in a body cavity, the handheld probe housing a probe opticalsystem.
 26. The system of claim 25 wherein the probe optical systemcomprises a plurality of optical collection paths to couple thatplurality of optical paths into separate collection optical fibers. 27.The system of claim 26 wherein the probe optical system comprises apolarizer and an aperture such that at least one polarized lightcomponent from the tissue is collected on one of the plurality ofoptical paths.
 28. The system of claim 1 further comprising a scanningsystem for scanning light across a region of interest to provide animage of the region.
 29. The system of claim 28 further comprisingsimultaneously scanning with a plurality of light spots.
 30. A methodspectroscopic imaging comprising: illuminating a material with a lightsource; detecting light with an imaging detector that detectsfluorescent and reflectance data to provide an image of the material;and processing the data with an data processor that processes the imageto characterize the material.
 31. The method of claim 30 furthercomprising providing a mask and plurality of filters.
 32. The method ofclaim 30 further comprising providing a moveable mirror device.
 33. Themethod of claim 30 further comprising providing a computer program todiagnose cancerous tissue.
 34. The method of claim 30 further comprisingusing a collimating light source.
 35. The method of claim 34 furthercomprising illuminating an excised tissue sample with a laser.
 36. Themethod of claim 30 further comprising illuminating the material with abroadband light source.
 37. The method of claim 30 further comprisingproviding a tunable light source.
 38. The method of claim 30 furthercomprising providing a digital micromirror array.
 39. The method ofclaim 30 further comprising switching between a first light source and asecond light source with a movable mirror.
 40. The method of claim 30further comprising providing a light source including an opticalparametric oscillator.
 41. The method of claim 30 further comprisingproviding a spatial gate.
 42. The method of claim 30 further comprisingproviding an angular gate.
 43. The method of claim 30 further comprisingscanning light across a region of interest to provide an image of theregion.
 44. The method of claim 43 further comprising simultaneouslyscanning with a plurality of light spots.
 45. A system for angulargating of light scattering spectroscopy comprising: illuminating amaterial with light from a light source; detecting light with a detectorthat detects light scattered from the material at different azimuthalangles to provide spectral data; and a data processor that processes thespectral data to characterize the material.
 46. The system of claim 45further comprising a mask.
 47. The system of claim 45 further comprisinga computer program to diagnose cancerous tissue.
 48. The system of claim45 wherein the light source comprises a broadband light source.
 49. Thesystem of claim 45 further comprising a first polarizer for anilluminating beam.
 50. The system of claim 49 further comprising asecond polarizer for a collected beam.
 51. The system of claim 45further comprising a first iris and a second iris.
 52. The system ofclaim 45 further comprising a long pass filter.